The methods and measuring equipment based on optical coherence domain reflectometry (OCDR) and optical coherence tomography (OCT) constitute the most widely-used solutions according to the known prior art for biometrically measuring and/or tomographically imaging eye structures.
In these coherence-optical methods, coherent light is, with the aid of an interferometer, used for distance-measuring and imaging purposes on reflective and scattering samples. On the human eye, the methods supply measurable signals when scanning the depth as a result of changes in the refractive index occurring at optical interfaces and as a result of volume scattering.
Optical coherence tomography is a very sensitive and quick method for interferometric imaging, which has found widespread use, particularly in the medical sector and in basic research. The images of eye structures based on OCT scans are often used in ophthalmology for diagnosis and accompanying therapy, and also for planning interventions and for selecting implants.
The basic principle of the OCT method, described e.g. in U.S. Pat. No. 5,321,501, is based on white-light interferometry and compares the run-time of a signal with the aid of an interferometer (usually a Michelson interferometer). In the process, the arm with a known optical path length (=reference arm) is used as a reference for the measurement arm, in which the sample is situated. The interference of the signals from both arms results in a pattern, from which it is possible to determine the scattering amplitudes as a function of the optical delay between the arms and hence a depth-dependent scattering profile, which, analogously to ultrasound technology, is referred to as an A-scan. In the multidimensional scanning methods, the beam is then guided transversely in one or two directions, as a result of which it is possible to record a two-dimensional B-scan or a three-dimensional volume tomogram. Here, the amplitude values of the individual A-scans are displayed using linear or logarithmic grayscale or false-color values. The technique of recording individual A-scans is also referred to as optical coherence domain reflectometry (OCDR), compared to which OCT realizes two- or three-dimensional imaging as a result of lateral scanning.
Two different basic types have prevailed in the OCT methods used in ophthalmology. In the first type, the length of the reference arm is changed and the intensity of the interference is recorded continuously in order to determine the measurement values, without this taking the spectrum into account. This method is referred to as a “time domain” method (U.S. Pat. No. 5,321,501 A). By contrast, in the other method, which is referred to as a “frequency domain” method, the spectrum is captured and the interference of the individual spectral components is evaluated in order to determine the measurement values.
It is for this reason that one refers firstly to the signal being in the time domain and secondly to the signal being in the frequency domain. The advantage of the “frequency domain” method lies in the simple and quick simultaneous measurement, in which it is possible to establish complete information in respect of the depth without requiring moveable parts. This increases the stability and the speed (U.S. Pat. No. 7,330,270 B2), as a result of which three-dimensional OCT recordings in particular were rendered possible.
In the case of frequency domain OCT, a distinction is furthermore made as to whether the spectral information is obtained by means of a spectrometer (“spectral domain OCT”, SD-OCT) or by means of spectral tuning of the light source (“swept source OCT”, SS-OCT).
A great technological advantage of OCT lies in the depth resolution being decoupled from the transverse resolution. As a result, it is possible to obtain very good axial resolutions, particularly also in the case of limited numerical apertures, such as for example in order to be able to examine retinal layers in the axial direction with high (<20 μm) and highest (<4 μm) resolutions, despite an aperture restriction by the pupil. The OCT measurement, which is based on back-scattering and reflection and hence is contactless, thus allows the generation of microscopic images of living tissue (in vivo). A further advantage lies in the efficient suppression of incoherent disturbance-light portions. In this case, the “axial direction” means the direction of the depth profile displayed in the A-scan. This direction may also vary in parts of the A-scan as a result of local refractions, but usually is almost parallel to the optical axis or to the visual axis from the corneal vertex to the fovea of an eye to be examined.
A first application of the display of the overall depth profile of the back-scattering from the eye, based on coherence-reflectometric measurements (OCDR), is described by F. Lexer et al. in [1]. Here, emphasis is once again placed on the fact that the precise knowledge of the intraocular distances is an important aid in modern ophthalmology, for example for fitting intraocular lens implants. While the axial eye length and the anterior chamber depth are mandatory for precise calculations of the refractions of intraocular lenses for cataract operations, the precise measurement of the corneal thickness is important for refractive surgery. Determining the thickness of the retinal layers can assist in diagnosing various disorders and monitoring the therapeutic effects. The approach to the solution described in [1] is based on an average-quality SS-FD-OCDR system and affords the possibility of measuring the distances of all optical areas in the eye over the whole eye length. Although it was still possible to obtain a high resolution when scanning the entire measurement region of model eyes using the approach to the solution, this was no longer possible during the simultaneous “in vivo” measurement of three intraocular distances. It was found that the approach to the solution for measuring intraocular distances with an accuracy of at best 30 μm achieves an adequate resolution; however, it is no longer suitable for high-resolution OCDR or OCT applications.
The laid-open applications US 2007/216909 A1, US 2007/291277 A1 and US 2008/100612 A1 describe SD-OCT systems which comprise a switchable focus and/or a switchable reference plane (zero-delay) of the OCT arrangement. Here, the focus should lie in the region of the retina in the case of a retinal scan and in the region of the cornea in the case of a cornea scan. Hence an OCT scan with a high lateral resolution of the anterior or posterior section of the eye is possible. Furthermore, a retinal scan with a scan rotation point in the iris/pupil plane is described as expedient. The solutions proposed here allow both high-resolution two-dimensional scans and three-dimensional scans (data cubes) to be recorded and evaluated.
A further OCT system based on the “frequency domain” is described by Walsh et al. in WO 2010/009447 A2. The spectral information is obtained either by use of a spectrometer (SD-OCT) or by use of a spectrally tunable light source (swept source, SS-OCT). Herein, the eye can be displayed in a plurality of sections along the optical axis or as a whole by application of an A-, B- or C-scan. The solution describes both a method for the whole eye scan and also partial scans arranged next to one another. The necessity of the scan rotation point being in the pupil for a retinal scan is also highlighted here. Furthermore, many ophthalmological application options for whole-eye scans are described.
The OCDR system described in DE 10 2008 051272 A1 serves for interferometrically measuring eye-section lengths over the whole eye length. A laterally scanning OCDR system is described, in which the focus too is variable or switchable in order to realize optimum A-scan signals by combination. No solution is proposed for displaying combined partial- or whole-eye OCT scans in an anatomically correct fashion. The radiation back-scattered from the interfaces of the eye is recorded by interferometry and a measurement signal displaying structures of the eye is created by time-domain, spectral-domain or Fourier-domain coherence reflectometry. This OCDR system provides a solution by which preferably a plurality of very precise partial distance measurements should simultaneously take place on the eye. The proposed OCDR system provides a solution by which eye section lengths can be measured in a very precise fashion over the entire eye length. Tomographic OCT recordings of the anterior and posterior eye sections by means of an A-, B- or C-scan are not possible using this system.
An OCDR system, which is based on long-coherent, tunable lasers (swept sources) and has over 40 mm scan depth in tissue, was proposed in DE 10 2008 063 225 A1. In particular, good signal ratios and low movement sensitivity on all interfaces of the eye can be realized well with this, even in the case of very long eyes, as is also shown in one example. An OCT system with a depth range of almost 35 mm, based on a relatively long-coherent, tunable source, is described by Ch. Chong et al. in [2]. In principle, the approach to the solution can realize tomographic images of the whole eye, in which images the contours of cornea, iris, lens and retina are rudimentarily visible. However, in the experimental implementation, it is difficult to see the details of the segments because implemented lateral resolutions and the signal strengths are rather low. As a result of the strong signal attenuation due to the insufficient coherence length of the source of only 28 mm, it was only possible to measure pig's eyes with a geometric length of approximately 20 mm. The system is inadequate for human eyes with a geometric length of up to 40 mm.
Furthermore, in [3] and [4], Reinstein et al. describe depth-resolved eye scans that can be used to display regions in the eye that cannot be displayed by means of OCT. By way of example, the prior art has disclosed highly-resolved ultrasound representations of anterior regions of the eye, including edge region of the lens of the eye behind the iris or the position of IOLs including the haptics behind the iris.
Registering OCT scans amongst one another or spatial referencing and correction by means of reference information from other, non-depth-resolving measurement systems, for example with height information of topographies, as described by Tang et al. in [3], is advantageous for the spatial assignment of measurement data. In addition to partial distance measurements, topographies or keratometries are required parameters for fitting refractive intraocular implants such as IOLs.
DE 43 09 056 describes another measurement method, based on short coherence, in which light from a broad-band light source is radiated into the sample and the light scattered back from various depths is analyzed spectrally. The depth information is obtained from a Fourier transform of the detected signal. This method is referred to as spectral domain OCDR (SD OCDR) or, since a Fourier transform is used, else as Fourier domain OCDR (FD OCDR). This category also includes the swept source OCDR (SS OCDR), as described in article [5] by S. H. Yun et al., in which the light source is spectrally tuned and the signal captured by the detector likewise contains the depth information after a Fourier transform. In this case, the imaging required for implementing optical coherence tomography (OCT), as already demonstrated in U.S. Pat. No. 5,321,501 for time domain OCT (TD OCT), is implemented by means of galvo scanners, which laterally deflect the measurement beam over the sample.
A first attempt at applying SS OCDR in optical biometry was described in [6] by F. Lexer, C. K. Hitzenberger, A. F. Fercher and M. Kulhavy. This solution showed that, in principle, it is possible to measure the intraocular distances within the eye; however, the measurement accuracy of 0.82 mm was much too inaccurate.
An improvement to this solution was disclosed by C. K. Hitzenberger, M. Kulhavy, F. Lexer, A. Baumgartner in [7]: “In-vivo intraocular ranging by wavelength tuning interferometry”, SPIE [3251-6] 1998. Here a resolution of 0.15 mm was achieved, but this still does not meet the requirements. However, in order to restrict the residual errors of the determined IOL refraction to 1/10 diopter, the measurement accuracy for the eye length must be less than 30 μm.
In particular, a problem for OCDR and OCT methods on moving samples, such as e.g. the human eye, is that the sample may move during the measurement which, as discussed by S. H. Yun et al. in [8]: (2004), OPTICS EXPRESS 2977, can significantly reduce and falsify the signals. Conventional approaches for rectifying the problem are complicated “tracking methods”, in which the movement of the sample is detected and the measurement beam position is updated.
By way of example, such approaches for compensating typical movements of a couple hundred micrometers per second are described by Hammer et al. in [9] (2005), Journal of Biomedical Optics 10(2), 024038 and in US 2006/105903. A disadvantage of such approaches is that despite the great technical complexity this always results in some updating errors as a result of the finite latency time of such systems, particularly in the case of very fast eye movements, e.g. saccades.
In general, coherence-optical systems are very sensitive and so secondary light portions, i.e. disturbance signals as a result of reflections and echoes at disturbance points, such as fiber couplers, fiber ends or else optical elements, cannot be suppressed by reflection-reducing measures to the extent that there are no disturbing interferences in the measurement region. Such disturbance signals are created by the reflection of portions of the measurement or reference light at elements or structures within the interferometric measurement arrangement and superpose on the actual measurement signals, making it significantly more difficult to evaluate the latter. By way of example, in this context fiber couplers, fiber ends or connection points of interferometric measurement arrangements embodied using fiber optics or else optical elements such as filters, lenses, beamsplitters, attenuators, circulators or the like should be considered as disturbance points.
It is known that disturbance signals based on such disturbance points occur if the optical path differences of the light portions in the interferometer that are reflected at the disturbance points and interfere with one another are smaller than the sought-after double measurement depth of the OCT system (or lens spacing is smaller than measurement depth). Since the sensitivity of an OCT system can easily be more than 90 or 100 dB, it is generally not possible to reduce the light portions generated at the disturbance points directly. By way of example, a common anti-reflective measure only brings about an attenuation of the reflection by 20 . . . 30 dB. Although ever-higher demands on the return losses of fiber couplers led to the development of plug-in connections in the form of so-called HRL (high return loss) or APC (angled physical contact) plugs, in which the angled, anti-reflection-coated fiber ends reflect in the region of −40 to −60 dB; however, compared to the OCT sensitivities this still does not provide a sufficient disturbance reflection suppression.
The effects of such disturbance signals are explained in more detail below on the basis of two examples.
In this respect, FIG. 1 shows disturbance signals, created by reflections, in an interferometric measurement arrangement. The light beam 102 emanating from the illumination source 101 is split into a measurement beam 104 and a reference beam 105 by a fiber coupler 103. In the case illustrated here, the aforementioned light beams are routed in optical fibers (single mode fibers), which, in the following text, is not mentioned individually in each case. Parts of the measurement beam 104 are reflected by the eye 109 and by the fiber end 112 (denoted by an x) and reach a further fiber coupler 106 via fiber coupler 103 and fiber 114. There the superposition on the reference beam 105 takes place and the resultant interference signals are routed to the two detectors 107 and 108. After amplifying the differences between the interference signals in the amplifier 110, a data processing unit 111 generates the measurement result, here in the form of an A-scan 115. The measurement result is influenced by the fiber-end reflection 112 if the optical path lengths L1 (from fiber coupler 103 to fiber coupler 106) and L2 (from fiber coupler 103 to the fiber end 112, back to the fiber coupler 103 and on to fiber coupler 106 via fiber 114) have a difference that is less than the sought-after double measurement region ZOCT. Under the condition that the magnitude of the path difference between the optical path lengths L1 and L2 is less than the sought-after double measurement region ZOCT, there is interference between the two light portions and hence the measurement result is influenced.
In the following text, ZOCT should be defined as the one-sided measurement depth in the direction of the measurement beam before or after the zero point of the interferometer. This is half the optical path length of the light cycle up to the measurement object and back. The zero point of the interferometer is the imagined point on the measurement beam where the optical path lengths of the reference arm and the measurement arm have the same length. An optical path length is the effective path length, taking into account the refractive indices of the media through which the light travels.
In the interferometric measurement arrangement as per FIG. 2, the light is split and superposed, and the interference signals are detected and evaluated, as above, but the disturbance signal is created at the connection point 113 (denoted by an x) of the fiber 105 between fiber coupler 103 and fiber coupler 106. In this case, the connection point 113 influences the measurement result if the optical path lengths L1 (from connection point 113 to fiber coupler 106) and L2 (from connection point 113 to fiber coupler 103 and on to fiber coupler 106 via fiber 114) have a difference that is smaller than the sought-after double measurement region ZOCT.
In order to prevent, or at least in order to minimize, the disturbance signals that are created at such disturbance points, the prior art only takes up partly effective, reflection-reducing measures, such as e.g. beamsplitters with oblique or anti-reflection-coated faces. Moreover, the cavities of the tunable light sources are selected such that measurement signals are only created in the sample arm and thus do not cause autocorrelation signals within the measurement region.
However, this can be avoided by suitably positioning the components and the reflection zones thereof, for example as per US 2007/0291277 A1. Then, influencing the measurement result is avoided if the following condition is satisfied:|L1−L2|>2*ZOCT  (1)
with L1: path length 1,                L2: path length 2, and        ZOCT: measurement region.        
Furthermore, the prior art has disclosed that the illumination source must be matched to the sought-after measurement region ZOCT of the interferometric arrangement in order to allow optimum signal generation and the evaluation thereof. In particular, care has to be taken when using laser or semiconductor light sources, e.g. superluminescent diodes (SLEDs) as well, that those light sources are selected whose cavity has an optical path length that is greater than the sought-after measurement region ZOCT in order to prevent certain autocorrelation signal artifacts in the measurement region.
On the other hand, such a signal is deliberately generated in dual-beam interferometers in order, for example, to be able to establish movement-independent eye lengths.
In this respect, FIG. 3a shows a sketch of the principle of an interferometric dual-beam measurement arrangement. The light beam 202 emanating from the illumination source 201 is split into two partial beams 204 and 205 by the beamsplitter 216. Partial beam 204 is delayed by the value LR by a reflection element 217. If the eye length AL now corresponds to the delay length LR, the delayed partial beam 204, after it was reflected on the cornea, interferes with the reflected, non-delayed partial beam 205 after the latter was reflected on the retina. A corresponding interference signal is created on the detector 207 and the data processing unit 211 generates the measurement result from said signal, here as an A-scan 115.
While a delayed light portion is used in the dual-beam interferometer for the measurement, the interferometric measurement arrangement as per FIG. 3b only wants measurement signals that result from the interference of reference light and light scattered back from the sample. Delayed light portions are unwanted in this example, although these can be emanated by the light source. In general, the delay of part of the light in coherent laser or semiconductor light sources is brought about by virtue of the fact that only some of the light cycling in the resonator is decoupled. The light portion that is routed back into the resonator at the decoupling point returns to the decoupling point after one cycle through the resonator. Thus, this produces light portions that are delayed by one or else more resonator cycles and are able to interfere with one another. In this example, the light beam 202 emanating from the light source 201 is split into a reference beam 204 and a measurement beam 205 at the fiber coupler 203. The light scattered back from the reflection element 217 and from the measurement object 209 is superposed in the fiber coupler 203. The created interference signals are measured by the spectrometer 218 and converted into the measurement result, in this case an A-scan 215, in the data processing unit 211. As described above using the dual-beam interferometer, delayed and non-delayed light portions can yield measurement signals if they impinge on structures with suitable spacing in the measurement object. Moreover, so-called autocorrelation signals may be created if the coherence length of the source is sufficiently large compared to the resonator length and, as a result thereof, light from different resonator cycles interferes with itself If the resonator length LR of the illumination source is now selected to be greater than the measurement region ZOCT thereof, it should be possible to assume that no disturbing interference signals are created.
Here, the resonator length LR should be understood to mean the optical path length that the light requires for one cycle in the illumination source. Hence, in the case of a ring laser, the resonator length LR corresponds to the length of the ring fiber. By contrast, in the case of a laser with a resonator, the resonator length LR is calculated from twice the length of the resonator because the light is only decoupled on one side and thus travels to and fro.
However, daily routine established that even if the two conditions are met, i.e. both the path differences of the disturbance-light portions created at a disturbance point and the cavity of the utilized light sources are selected to be greater than the sought-after measurement region ZOCT, disturbance signals which adversely affect the evaluation of the measurement may, under certain circumstances, be created in the interferometric measurement arrangement. Here, this problem occurred particularly in the case of long-coherent swept-source-based OCDR systems and/or in the case of a large measurement region.